Magnetic resonance imaging apparatus

ABSTRACT

A magnetic resonance imaging apparatus according to an embodiment includes an imaging unit configured to carry out magnetic resonance imaging of a patient using a transmitting QD coil which allows at least one of phase and amplitude of a radio-frequency transmit pulse on at least one input channel of the transmitting QD coil to be adjusted independently of each other, and an adjustment unit configured to adjust at least one of the phase and the amplitude of the radio-frequency transmit pulse according to imaging conditions.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is based upon and claims the benefit of priority fromJapanese Patent Applications No. 2009-269210 filed on Nov. 26, 2009, andNo. 2010-256379 filed on Nov. 17, 2010, the entire contents of which areincorporated herein by reference.

FIELD

Embodiments described herein relate generally to a Magnetic ResonanceImaging apparatus.

BACKGROUND

Magnetic resonance imaging is an imaging method which magneticallyexcites nuclear spins of a patient placed in a static magnetic fieldwith an RF signal at the Larmor frequency to reconstruct an image usingan NMR (nuclear magnetic resonance) signal resulting from theexcitation.

In the magnetic resonance imaging, when frequency of a transmitted RFpulse increases, an RF magnetic field in the patient becomesnon-uniform. The non-uniformity of the RF magnetic field is alsoreferred to as B1 non-uniformity. Thus, it is important to correct theB1 non-uniformity.

For example, a B1 correction method has been proposed which determinesamplitude and phase of an RF transmit pulse based on aspect ratio. Also,a method is known which collects a B1 map by changing the amplitude andphase of an RF transmit pulse and thereby determines the amplitude andphase of the RF transmit pulse.

However, an MRI apparatus with high magnetic field strength have come tobe developed recently. Consequently, the frequency of RF pulses hasbecome greater, making it difficult to resolve the B1 non-uniformityeven if an RF pulse is transmitted using the amplitude and phasedetermined based on aspect ratio as is conventionally the case. On theother hand, when collecting a B1 map by changing the amplitude and phaseof an RF pulse as is conventionally the case, there are problems ofincreased imaging time and difficulty to determine B1 non-uniformity.

Thus, there is demand for a magnetic resonance imaging apparatus whichcan further improve B1 uniformity in a simple way.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a configuration diagram showing an embodiment of a magneticresonance imaging apparatus according to the present invention;

FIG. 2 is a diagram showing a detailed configuration example of atransmitting QD coil and transmitter;

FIG. 3 is a functional block diagram of a computer;

FIG. 4 is a diagram showing an example of a phase-amplitude look uptable;

FIG. 5 is a flowchart showing exemplary procedures for adjusting phaseand amplitude of an RF transmit pulse to improve B1 uniformity beforeimaging;

FIG. 6 is a diagram showing a DSE sequence for collecting data todetermine an optimal phase of an RF transmit pulse;

FIG. 7 is a diagram illustrating a method for determining an optimalphase Φ2opt from a curve which represents changes in non-uniformity ofimage values vs. changes in phase Φ2 of an RF transmit pulse;

FIG. 8 is a diagram illustrating a method for determining the optimalphase Φ2opt from uniformity of image values and forward power of arefocusing pulse;

FIGS. 9A and 98 are diagrams illustrating a method for determining theoptimal phase Φ2opt of an RF transmit pulse using symmetry of B1 mapdata in the head;

FIG. 10 is a flowchart showing exemplary procedures for adjusting thephase Φ2 and gain G2 based on two sets of B1 map data collected by meansof independent transmission via two channels ch1 and ch2;

FIGS. 11A and 118 are diagrams showing an example of two sets of B1 mapdata collected by using the two channels ch1 and ch2 separately.

FIGS. 12A and 128 are diagrams each showing an example which compares B1map data collected before and after correction of the phase Φ2, wherethe phase Φ2 is corrected by being shifted 60° from its initial value;and

FIGS. 13A and 13B are diagrams each showing an example which comparestwo sets of B1 map data collected by correcting the gain G2 by differentamounts of correction Δgain2, where the amounts of correction Δgain2used are −5 dB and +5 dB of an initial value of the gain G2.

DETAILED DESCRIPTION

An embodiment of a magnetic resonance imaging apparatus will bedescribed with reference to the accompanying drawings. The magneticresonance imaging apparatus according to the embodiment includes animaging unit configured to carry out magnetic resonance imaging of apatient using a transmitting QD coil which allows at least one of phaseand amplitude of a radio-frequency transmit pulse on at least one inputchannel of the transmitting QD coil to be adjusted independently of eachother, and an adjustment unit configured to adjust at least one of thephase and the amplitude of the radio-frequency transmit pulse accordingto imaging conditions.

(Configuration and Functionality)

FIG. 1 is a configuration diagram showing an embodiment of a magneticresonance imaging apparatus according to the present invention.

The magnetic resonance imaging apparatus 20 includes a static magneticfield magnet 21 which is cylindrical in shape and configured to generatea static magnetic field, a shim coil 22 installed in the static magneticfield magnet 21, a gradient coil 23, and RF coils 24.

Also, the magnetic resonance imaging apparatus 20 includes a controlsystem 25. The control system 25 has a static magnetic field powersupply 26, a gradient power supply 27, a shim coil power supply 28, atransmitter 29, a receiver 30, a sequence controller 31, and a computer32. The gradient power supply 27 of the control system 25 includes anX-axis gradient power supply 27 x, Y-axis gradient power supply 27 y,and Z-axis gradient power supply 27 z. Also, the computer 32 is equippedwith an input device 33, a display device 34, a processing device 35,and a storage device 36.

The static magnetic field magnet 21, which is connected with the staticmagnetic field power supply 26, has a capability to form a staticmagnetic field in an imaging area using an electric current suppliedfrom the static magnetic field power supply 26. The static magneticfield magnet 21 is often made of a superconducting coil and is connectedwith the static magnetic field power supply 26 during excitation to drawelectric current, but generally becomes disconnected once excited. Onthe other hand, there are cases in which the static magnetic fieldmagnet 21 is made of a permanent magnet without installation of thestatic magnetic field power supply 26.

The shim coil 22 which is cylindrical in shape is installed coaxiallyinside the static magnetic field magnet 21. The shim coil 22 isconnected with the shim coil power supply 28, supplied with electriccurrent from the shim coil power supply 28, and configured to make thestatic magnetic field uniform.

The gradient coil 23 is made up of an X-axis gradient coil 23 x, Y-axisgradient coil 23 y, and Z-axis gradient coil 23 z and formed into acylindrical shape inside the static magnetic field magnet 21. A bed 37is installed in the gradient coil 23 and used as an imaging area with apatient P placed thereon. The RF coils 24 include a whole body coil(WBC) incorporated in a gantry and used to transmit and receive RFsignals, and a local coil installed near the bed 37 or patient P andused to receive RF signals.

The gradient coil 23 is connected with the gradient power supply 27. TheX-axis gradient coil 23 x, Y-axis gradient coil 23 y, and Z-axisgradient coil 23 z of the gradient coil 23 are connected, respectively,with the X-axis gradient power supply 27 x, Y-axis gradient power supply27 y, and Z-axis gradient power supply 27 z of the gradient power supply27.

A gradient magnetic field Gx in an X-axis direction, gradient magneticfield Gy in a Y-axis direction, and gradient magnetic field Gz in aZ-axis direction are formed, respectively, in the imaging area byelectric currents supplied to the X-axis gradient coil 23 x, Y-axisgradient coil 23 y, and Z-axis gradient coil 23 z, respectively, fromthe X-axis gradient power supply 27 x, Y-axis gradient power supply 27y, and Z-axis gradient power supply 27 z.

The RF coils 24 are connected to the transmitter 29 and/or receiver 30.Transmitting RF coils 24 have a capability to receive an RF signal fromthe transmitter 29 and transmit the RF signal to the patient P.Receiving RF coils 24 have a capability to receive an NMR signalgenerated when nuclear spins in the patient P is excited by the RFsignal and supply the NMR signal to the receiver 30.

FIG. 2 is a diagram showing a detailed configuration example of thetransmitting RF coil 24 and transmitter 29 shown in FIG. 1.

The whole body coil 24A which is a transmitting RF coil 24 is configuredas a QD coil. The QD coil is normally fed with transmit signals from twotransmission channels ch1 and ch2 and generates magnetic fieldsspatially orthogonal to each other using the first transmission channelch1 and second transmission channel ch2. Available types of QD coilinclude a cylindrical bird-cage type or a type made of independent twocoils. With the bird-cage type QD coil, transmit signals through thechannels ch1 and ch2 are fed, respectively, into two feeding points 90degrees apart from each other on the cylinder. On the other hand, withthe independent two coil type, transmit signals from the channels ch1and ch2 are fed separately into two coils, respectively.

At RF transmit pulse outputted from an RF amplifier 29A is split by a90-degree splitter 209 into two signals which are 90 degreeselectrically out of phase with each other. These two signals are fed tothe first and second transmission channels ch1 and ch2, respectively.

The transmitting QD coil applies two RF signals 90 degrees electricallyout of phase with each other in directions spatially orthogonal to eachother and thereby forms a perfectly uniform magnetic field on an XYplane of an imaging area under ideal conditions.

Actually, however, such ideal conditions do not exist. For example, theuniformity of the magnetic field is affected greatly by positionalrelationship between the position of the transmitting QD coil and animaging region (the head, abdomen, or wrists of the patient). That is,even if a transmitting QD coil (e.g., whole body coil) is used as atransmitting coil, the uniformity of the magnetic field is notnecessarily ensured depending on the imaging region, and non-uniformstate varies with the imaging region.

Also, the uniformity of the magnetic field is affected greatly by thetype of coil used. For example, even if a transmitting QD coil is used,the uniformity of the magnetic field varies depending on what receivercoil is used: a torso coil or wrist coil.

Thus, as shown in FIG. 2, phase adjustment units 210A and 210B as wellas gain adjustment units 211A and 211B are installed on the first andsecond transmission channels ch1 and ch2 as means for adjusting theuniformity of the magnetic field.

The phase adjustment units 210A and 210B and the gain adjustment units211A and 2118 allow adjustment of relative phase (phase balance) andrelative gain (gain balance) between the channels. The phase adjustmentunits 210A and 210B as well as the gain adjustment units 211A and 2118may be installed separately on the respective channels as shown in FIG.2 or installed only on one of the channels.

In the following description, it is assumed that the phase balance andgain balance between the channels are adjusted by varying phase Φ2 andgain G2 of the second transmission channel ch2, with phase Φ1 and gainG1 of the first transmission channel ch1 fixed at 0 degrees and 0 dB,respectively. Although variable ranges of the phase Φ2 and gain G2 arenot limited particularly, the phase Φ2 may be set, for example, between−90° and 90° and the gain G2 may be set, for example, between −10 dB and10 dB.

Incidentally, only one of the phase Φ2 and gain G2 may be madeadjustable.

The first and second transmission channels ch1 and ch2 are provided witha capability to measure power of a forward pulse of the RF transmitpulse to be outputted to the QD coil and a capability to measure powerof reflected pulses coming from the QD coil.

Returning to FIG. 1, the sequence controller 31 of the control system 25is connected with the gradient power supply 27, transmitter 29, andreceiver 30. The sequence controller 31 has a capability to storecontrol information such as sequence information needed to drive thegradient power supply 27, transmitter 29, and receiver 30, where thesequence information describes operational control information such asintensity, application duration, and application timing of a pulsedcurrent to be applied to the gradient power supply 27. The sequencecontroller 31 also has a capability to drive the gradient power supply27, transmitter 29, and receiver 30 according to a stored predeterminedsequence and thereby generate an X-axis gradient magnetic field Gx,Y-axis gradient magnetic field Gy, Z-axis gradient magnetic field Gz,and RF signal.

The sequence controller 31 is configured to receive raw data and supplythe raw data to the computer 32, where the raw data is complex dataproduced by the receiver 30 through detection and A/D (analog todigital) conversion of NMR signals.

The transmitter 29 supplies an RF signal to the RF coil 24 based on thecontrol information received from the sequence controller 31. On theother hand, the receiver 30 detects the NMR signal received from the RFcoil 24, performs necessary signal processing and A/D conversion, andthereby generates raw data which is digitized complex data, and suppliesthe generated raw data to the sequence controller 31.

The computer 32 can implement various functions by causing theprocessing device 35 to execute programs stored in the storage device 36of the computer 32. However, specific circuits with various functionsmay be installed in the magnetic resonance imaging apparatus 20 withoutdepending on programs.

FIG. 3 is a functional block diagram of the computer 32 shown in FIG. 1.

The computer 32 functions as an imaging condition setting unit 40,sequence controller control unit 41, and data processing unit 42 basedon programs. The imaging condition setting unit 40 includes aphase-amplitude LUT (look up table) 40A and a phase and amplitudeadjustment unit 40B.

The imaging condition setting unit 40 has a capability to set imagingconditions including an imaging region, coil type, and pulse sequencebased on command information from the input device 33, and supplyvarious parameters based on the set imaging conditions to the sequencecontroller control unit 41.

The phase-amplitude LUT 40A prestores appropriate phase and amplitudevalues of the RF transmit pulse, which can be used as reference. In thephase-amplitude LUT 40A, the phase and amplitude values are classifiedaccording to imaging conditions such as the imaging region or the typeof transmitter coil or receiver coil used.

FIG. 4 is a diagram showing an example of the phase-amplitude LUT 40A.The phase-amplitude LUT 40A stores values of the phase Φ2 and gain G2 ofthe second transmission channel ch2 in terms of differences from thefirst transmission channel ch1. The values of the phase Φ2 and gain G2in the phase-amplitude LUT 40A are associated with imaging regions suchas a head, abdomen, or wrists of the patient, as well as coil types suchas a phased array head coil, T/R (transmit/receive) head coil, torsocoil, or wrist coil.

The phase-amplitude LUT 40A also stores gain correction values (ΔG)according to body weight.

The phase Φ2 and gain G2 of the RF transmit pulse in the phase-amplitudeLUT 40A can be determined by analysis of data collected by past scans ortest scans, or through simulations. Alternatively, the phase Φ2 and gainG2 in the phase-amplitude LUT 40A may be determined by a method similarto the one used for a readjustment sequence described later.

Prior to an imaging scan, the phase and amplitude adjustment unit 40Badjusts the phase Φ2 and gain G2 of the RF transmit pulse to beappropriate values according to the imaging conditions by referring tothe phase-amplitude LUT 40A. In addition, the phase and amplitudeadjustment unit 40B has a capability to readjust the adjusted phase Φ2and gain G2 of the RF transmit pulse further, based on image data andthe like acquired from the sequence controller control unit 41, and setthe finally-established phase Φ2 and gain G2 of the RF transmit pulse asimaging scan parameters on the sequence controller control unit 41.

The sequence controller control unit 41 outputs various parametersincluding a pulse sequence to the sequence controller 31 based oninformation from the input device 33 and imaging condition setting unit40. Also, the sequence controller control unit 41 has a capability toreceive raw data from the sequence controller 31 and supply the raw dataas k-space data to the data processing unit 42.

The data processing unit 42 has a capability to apply an imagereconstruction process including a Fourier transform (FT) to the k-spacedata, and thereby generate image data as well as a capability to applynecessary image processing to the image data and display the resultingimage data on the display device 34.

(Operation)

Next, operation of the magnetic resonance imaging apparatus 20 will bedescribed. In the example described below, it is assumed that the phaseΦ2 and gain G2 of the second transmission channel ch2 are variable whilethe phase Φ1 and gain G1 of the first transmission channel ch1 arefixed.

FIG. 5 is a flowchart showing exemplary procedures for adjusting thephase and amplitude of an RF transmit pulse to improve B1 uniformitybefore imaging on the magnetic resonance imaging apparatus 20 shown inFIG. 1.

To begin with, the patient P is placed on the bed 37 in advance and astatic magnetic field is formed in an imaging area of the staticmagnetic field magnet 21 (superconducting magnet) excited by the staticmagnetic field power supply 26. Also, an electric current is supplied tothe shim coil 22 from the shim coil power supply 28, and consequentlythe static magnetic field formed in the imaging area is made uniform.

In Step S1, the phase Φ2 and gain G2 of the RF transmit pulsetransmitted from the second transmission channel ch2 are set toappropriate values according to the imaging conditions such as theimaging region of the patient and the type of coils used for imaging.

Specifically, the phase and amplitude adjustment unit 40B of the imagingcondition setting unit 40 acquires the phase Φ2 and gain G2 of the RFtransmit pulse which meet the imaging conditions by referring to thephase-amplitude LUT 40A and outputs the acquired phase Φ2 and gain G2 tothe sequence controller control unit 41 so that the phase Φ2 and gain G2will be set on the phase adjustment unit 210B and gain adjustment unit211B of the transmitter 29, respectively.

As imaging conditions, the body weight and lying position of the patientmay be used in addition to the imaging region of the patient and thetype of coils used for imaging. For example, as shown in FIG. 3, whenthe body weight of the patient exceeds 80 kg, a gain correction value ΔGis added to the gain G2.

Also, when the lying position of the patient is lateral recumbentposition (which results when the patient turns 90 degrees sideways froma supine position), the phase Φ2 and gain G2 of the RF transmit pulseacquired from the phase-amplitude LUT 40A is multiplied by −1. Forexample, if the phase Φ2 and gain G2 acquired from the phase-amplitudeLUT 40A are 140 degrees and 3 dB, respectively, the phase Φ2 and gain G2are corrected by being multiplied by −1 and consequently the values ofthe phase Φ2 and gain G2 read out of the phase-amplitude LOT 40A arereplaced by −140 degrees and −3 dB, respectively.

Although the phase Φ2 and gain G2 of the RF transmit pulse can be set toappropriate values according to the imaging conditions by referring tothe phase-amplitude LUT 40A, they can be set to more appropriate valuesby readjustment procedures in Step S2 and subsequent steps. In theprocesses of Step S2 and subsequent steps, the values of the phase Φ2and gain G2 stored in the phase-amplitude LUT 40A are set as initialvalues in the phase adjustment unit 210B and gain adjustment unit 211Bof the transmitter 29 and the phase Φ2 and gain G2 are readjusted basedon image data acquired while changing the phase Φ2 and gain G2 from theinitial values. The readjustment is performed as follows. First, in StepS2, an optimal output level RFlevel of the RF transmit pulse isdetermined. Specifically, signals are collected while changing an outputpower level of the RF transmit pulse using a known data collectionmethod such as an SE (spin echo) method, and then the output levelRFlevel which maximizes strength of the collected signals is determined.

Next, in Step S3, a target area is set to collect data for use toreadjust the phase Φ2 and gain G2 of the RF transmit pulse. The targetarea for readjusting is set in correspondence to an imaging area for theimaging scan. An axial section is often selected as a target area, but alocal area such as a sagittal section or coronal section may beselected.

Next, in Step S4, an optimal phase Φ2opt of the RF transmit pulse isdetermined based on the data collected while changing the phase Φ2 ofthe RF transmit pulse.

Specifically, signals are collected by changing the phase Φ2 of the RFtransmit pulse without phase encoding. For example, if the initial valueof the phase Φ2 of the RF transmit pulse stored in the phase-amplitudeLUT 40A is 30°, the signals are collected by changing the phase Φ2 ofthe RF transmit pulse at intervals of 20° such as −90°, −70°, −50°, . .. , 90°. Then, the optimal phase Φ2opt which provides the highest B1uniformity and consequently maximizes the strength of the signalscollected from the target area is calculated. This is equivalent todetermining such a phase Φ2 angle of the second channel ch2 that willcause the B1 transmitted from the first channel ch1 and the B1transmitted from the second channel ch2 to intersect each other at rightangles.

In so doing, a single-axis or two-axis slice selection SE method or asingle-axis, two-axis or three-axis slice selection DSE (Double SpinEcho) method can be used as a data collection method.

FIG. 6 is a diagram showing a DSE sequence for collecting data todetermine the optimal phase Φ2opt of the RF transmit pulse.

In FIG. 6, reference character RE denotes the RF transmit pulse, Gssdenotes a gradient magnetic field in a slice selection direction, Grodenotes a gradient magnetic field in a readout direction, Gpe denotes agradient magnetic field in a phase encoding direction, and ECHO denotesecho signals collected.

As shown in FIG. 6, in the DSE sequence, an excitation pulse with a flipangle of α, the first refocusing pulse with a flip angle of β, and thesecond refocusing pulse with a flip angle of γ are applied, accompaniedby slice selections in the slice selection direction, readout direction,and phase encoding direction, respectively. A gradient pulse for readoutis applied after the application of the first refocusing pulse withoutapplication of a gradient pulse for phase encoding, and consequently thefirst echo signal is collected TE/2 (TE stands for echo time) after theexcitation pulse. Also, a gradient pulse for readout is applied afterthe application of the second refocusing pulse without application of agradient pulse for phase encoding, and consequently the second echosignal is collected TE/2 after the first echo signal.

If attenuation is ignored, strength Sse of the echo signal collected bythe SE method is given by Equation (1), where α is the flip angle of theexcitation pulse and β is the flip angle of the first refocusing pulse.Also, if attenuation is ignored, strength Sdse of the echo signalcollected second by the DSE method is given by Equation (2), where α isthe flip angle of the excitation pulse, β is the flip angle of the firstrefocusing pulse, and γ is the flip angle of the second refocusing pulseas described above.

Sse=sin(α)·sin²(β/2)  (1)

Sdse=sin(α)·sin²(β/2)·sin²(γ/2)  (2)

As can be seen from Equations (1) and (2), when the flip angle α isclose to 90° or the refocusing angle β is close to 180°, which are bothtypical values, the strengths Sse and Sdse change gradually, and thechanges in the signal strength are small compared to variations in a B1magnetic field. On the other hand, if the flip angle α and therefocusing angles β and γ in the SE or DSE method are set smaller thantypical flip angles, slopes of changes in the strengths Sse and Sdse areincreased. Consequently, a degree of the B1 non-uniformity can bereflected in the signals with high sensitivity. In other words, if theflip angles are set to be small, signals which are sensitive to thechanges in the B1 magnetic field can be collected, making it possible tomore accurately determine the phase Φ2 for increasing the B1 uniformity.

For example, in the case of the SE method, preferably the flip angles αand β of the excitation pulse and first refocusing pulse are set to 30°and −60°, respectively, rather than typical angles of 90° and −180°. Inthe case of the DSE method, preferably the flip angle α of theexcitation pulse and the flip angles β and γ of the refocusing pulsesare set to 30°, −60°, and −60° rather than typical angles of 90°, −180°,and −180°.

Based on the echo signals collected by the SE method or DSE method, anindex which represents non-uniformity of the image is calculated foreach value of the phase Φ2. The index which represents thenon-uniformity of the image may be an area where intensity of the imagealong one axis deviates from an average of the intensity, or a variancearound the average. The area or the variance may be obtained within arange in which deviation from an average exceeds a threshold (e.g., −20%and below and/or 20% and above). Specifically, an index S whichrepresents the non-uniformity of the image in the readout direction canbe calculated for each value of the phase Φ2 by computing 1D FFT(one-dimensional fast Fourier transformation) of the echo signalscollected at a small flip angles α and β of the excitation pulse andrefocusing pulse using the SE method or the second echo signal collectedat a small flip angles α, β, and γ of the excitation pulse andrefocusing pulses using the DSE method.

FIG. 7 is a diagram showing a curve of changes in the non-uniformityindex S of image values with respect to the phase Φ2 of the RF transmitpulse.

First, an average of the intensity of the image in the readout directionis calculated, then, the area or the variance is calculated as indexvalues S01, S02, S03, and so forth. Then, a known approximation methodsuch as polynomial fitting is applied to the index values S01, S02, S03,and so forth to obtain a non-uniformity curve such as represented by asolid line in FIG. 7. This makes it possible to determine the optimalphase Φ2opt which minimizes the non-uniformity curve.

In addition to the example described above, various methods areavailable for use to find the optimal phase Φ2opt. For example, with theDSE method, echo signals may be collected by setting only the flip angleγ of the second refocusing pulse to a value smaller than usual, with theflip angle α of the excitation pulse and flip angle β of the firstrefocusing pulse set to typical values of 90° and 180°, respectively. Inthis case, the phase Φ2 which minimizes the quotient obtained bydividing the non-uniformity index value of the image of the second echosignal by the non-uniformity index value of the image of the first echosignal may be used as the optimal phase Φ2opt. This approach providesthe advantage of being robust to the effects of body movements.

The optimal phase Φ2opt may also be determined based on a phaseΦ2-dependent non-uniformity index which involves three or more echosignals, by collecting the echo signals using a sequence for collectingthree or more echo signals. For example, a 2D image can be obtained bycollecting eight or so echo signals by single-shot or multi-shot(two-shot or so) imaging while performing phase encoding using the FSE(fast spin echo) method and then by applying a 2D (two-dimensional) FFTto the collected echo signals. Then, a non-uniformity index curve withrespect to the phase Φ2 can be created using pixel values of the 2Dimage data.

Another example involves collecting 1D or 2D B1 map data in differentphases Φ2 by a known technique which uses a gradient echo method andcreating a non-uniformity index curve with respect to the phase Φ2 basedon a representative value of the B1 map data for each phase Φ2. In thiscase, sensitivity to non-uniformity can be improved if intensity changesof the map data is enhanced by squaring or cubing the B1 map data beforecreating the non-uniformity index curve. For example, to obtain B1 mapdata only in the readout direction using a sequence without phaseencoding, a square value or cubic value of B1 map data can be calculatedfor each value of the phase Φ2 only in the readout direction, and anon-uniformity index curve with respect to the phase Φ2 can be created.

The optimal phase Φ2opt of the RF transmit pulse may be determined bytaking forward power FP of the refocusing pulse into consideration inaddition to the B1 non-uniformity index S.

Specifically, as shown in FIG. 8, among the phases Φ2 in which thenon-uniformity index S falls within a predetermined reference range(e.g., within ±10% of a minimum value of the non-uniformity index S), aphase Φ2 which minimizes a total value FP of forward powers FP1 and FP2of the refocusing pulse transmitted from the two transmission channelsch1 and ch2 may be designated as the optimal phase Φ2opt. The forwardpowers FP1 and FP2 of the refocusing pulse can be measured using asystem shown in FIG. 2 while the non-uniformity index S is measured, bychanging the phase Φ2. With this method, the phase Φ2 which providesgood B1 uniformity and reduces the forward power of the refocusing pulseis determined as the optimal phase Φ2opt.

Incidentally, if data collection time is likely to be long due toincrease of the number of conditions, the data collection time can bereduced by limiting the variable range of the phase Φ2, for example, tobetween 0° and 90°.

So far, description has been given of a technique for determining theoptimal phase Φ2opt from the B1 non-uniformity index S or B1 map dataacquired by changing the phase Φ2, however, the optimal phase Φ2opt canalso be found directly from a single set of B1 map data obtained with aphase Φ2int set as an initial value without changing the phase Φ2. FIGS.9A and 98 are diagrams illustrating a concept of this technique.

FIG. 9A is a diagram showing B1 map data of a head of the patientsuperimposed with a white line. The white line is obtained by linkingthe same values in the B1 map data and fitting a resulting contour to anelliptical model.

Generally, B1 map data of the head has a contour of a relatively stableshape. When the phases Φ2 of the first and second transmission channelsch1 and ch2 are in ideal conditions, an axial section of the head of thepatient lying in a supine position is longer in a horizontal directionthan in a vertical direction. Therefore, when the phase Φ2 is in idealconditions, a major axis of an ellipse to which the contour of the headis fitted, as shown by a solid line in FIG. 9B, is oriented in thehorizontal direction.

On the other hand, if the phase Φ2 deviates from the ideal conditions,the major axis of the ellipse to which the head contour is fitted, asshown by a broken line in FIG. 9B, has an angle Φ2 c with respect to thehorizontal direction rather than being horizontal. The angle in degreecorresponds to a deviation from an ideal value of phase Φ2 in degree.Therefore, by correcting the initial value Φ2int (deg) of the phase Φ2with which sloped B1 map data by the angle equal to the slope Φ2 c, theideal value of the phase Φ2, i.e., the optimal phase Φ2opt (deg), can beobtained.

Thus, with the present technique, the contour of the B1 map dataobtained in a phase Φ2int which is set as an initial value is fitted toan ellipse, and the deviation of the major axis of the resulting ellipsefrom the horizontal direction is used as a correction value Φ2 c of thephase Φ2. Then, the initial phase Φ2int (deg) is corrected by thecorrection value Φ2 c (using addition or subtraction) to obtain theoptimal phase Φ2opt. Although the present technique has limitedapplicability, being applicable only to regions, such as the head, whichhave an easily identifiable shape, the present technique can reduce thetime required for correction because the optimal phase Φ2opt can befound directly from a single set of B1 map data.

Next, in Step S5, optimal gain G2opt of the RE transmit pulse isdetermined based on data collected while changing the gain G2 of the RFtransmit pulse. In Step S5, the phase Φ2 of the RF transmit pulseremains fixed at the determined optimal phase Φ2opt.

Specifically, signals are collected by changing the gain G2 of the RFtransmit pulse. For example, if 2 dB is stored in the phase-amplitudeLUT 40A as an initial value of the gain G2 of the RF transmit pulse,signals are collected by changing the gain G2 of the RF transmit pulseat intervals of 2 dB such as −8 dB, −6 dB, −4 dB, . . . , 8 dB.Consequently, the optimal gain G2opt which provides high B1 uniformityand maximizes the strength of the signals collected from the target areais calculated. Incidentally, to compare multiple signals at differentgains G2, the output level RFlevel of the RE transmit pulse is adjustedso as to make total power of the transmit pulse outputted from the twochannels ch1 and ch2 constant. For example, if the output level RFlevelin initial state is 50 dB, the RF transmit pulse is outputted from thefirst channel ch1 at 50 dB. On the other hand, the RF transmit pulse isoutputted from the second channel ch2 at 52 dB, which is equal to theoutput level RFlevel of 50 dB plus the initial value of the gain G2 of 2dB. That is, the total power of the RF transmit pulse is 102 dB indecibel terms. On the other hand, if the gain G2 of the second channelch2 is set, for example, to +6 dB, the output level RFlevel is set to 48dB. In this case, the RF transmit pulses are outputted at 54 dB from thesecond channel ch2, and at 48 dB from the first channel ch1, for a totalpower of 102 dB all the same, thereby making the total power constant.

The optimal gain G2opt can also be determined from a B1 non-uniformityindex curve or B1 map data plotted or obtained at different gains G2, asin the case of the optimal phase Φ2opt. Also, in determining the optimalgain G2opt, the gain G2 which minimizes the forward power of therefocusing pulse can be selected out of the gains G2 at whichnon-uniformity falls within a predetermined reference range anddesignated as the optimal gain G2opt.

Incidentally, if data collection time is likely to be long due toincrease of the number of conditions, the data collection time can bereduced by limiting the variable range of the gain G2, for example, tobetween 0 dB and 6 dB.

Next, in Step S6, the output level RFlevel of the RF transmit pulse isreadjusted as required. That is, since the phase Φ2 and gain G2 of theRF transmit pulse have been readjusted to be the optimal phase Φ2opt andgain G2opt, the conditions for determining the optimal output levelRFlevel of the RF transmit pulse have been changed as well. Thus,desirably, the output level RFlevel of the RF transmit pulse isreadjusted under condition of the optimal phase Φ2opt and gain G2opt ofthe RF transmit pulse.

The readjusting of the output level RFlevel can be done in the samemanner as in Step S2, but another method may be used.

For example, echo signals can be collected from the target area whilechanging the output level RFlevel using the SE method or DSE method andthe output level RFlevel which maximizes the strength of the collectedecho signals can be used as a readjusted value. The target area is setin correspondence to the imaging area for the imaging scan.

Alternatively, B1 map data can be collected from the target area whilechanging the output level RFlevel using the SE method or DSE method andthe optimal value of the output level. RFlevel can be found based on theB1 map data obtained at each output level RFlevel. That is, a B1non-uniformity index curve can be created by plotting at each outputlevel RFlevel based on the variance and average value of the B1 map dataand the output level RFlevel which minimizes the B1 non-uniformity indexcan be used as a readjusted value. Alternatively, the output levelRFlevel corresponding to optimal B1 map data may be used as a readjustedvalue.

Still another example involves storing amounts of correction to adefault value of the output level RFlevel in the phase-amplitude LUT 40Aby classifying the amounts of correction according to imaging conditionssuch as the imaging region and the type of transmitter coil as well asto the phase Φ2 and gain G2 of the RF transmit pulse. Then, the phaseand amplitude adjustment unit 40B can acquire the appropriate amount ofcorrection according to imaging conditions, phase Φ2, and gain G2 fromthe phase-amplitude LUT 40A and readjust the output level RFlevelaccording to the acquired amount of correction.

In this way, the phase and amplitude adjustment unit 40B can find theoptimal phase Φ2opt, gain G2opt, and output level RFlevel of the RFtransmit pulse according to the imaging conditions. Incidentally, toreduce the data collection time, any of the steps other than Step S1 maybe omitted. Also, the optimal gain G2opt may be determined before theoptimal phase Φ2opt is determined with the gain G2 fixed. However, thephase Φ2 has a more dominant impact on the B1 uniformity than does thegain G2. Therefore, it is more desirable to determine the optimal phaseΦ2opt before determining the optimal gain G2opt.

Furthermore, to reduce the data collection time, echo signals can becollected by changing both phase Φ2 and gain G2 and then both optimalphase Φ2opt and optimal gain G2opt can be found at the same time. Thatis, the B1 non-uniformity index or B1 map data may be acquired bychanging both phase Φ2 and gain G2. Also, the forward power of therefocusing pulse may be acquired by changing both phase Φ2 and gain G2.Also, the forward power of the refocusing pulse can be obtained bychanging both the phase Φ2 and the gain G2. Then, both optimal phaseΦ2opt and optimal gain G2opt can be found simultaneously to furtherreduce the B1 non-uniformity and the forward power of the refocusingpulse.

With the technique described above, it is assumed that the proceduresfor readjustment of the phase Φ2 and gain G2 in Steps S2 to S6(especially Steps S3 to S5) are carried out. Therefore, the initialvalues of the phase Φ2 and gain G2 stored in the phase-amplitude LUT 40Ado not need to be very accurate.

Conversely, the procedures for readjustment of the phase Φ2 and gain G2in Steps S3 to S5 may be carried out in advance for each imaging regionand coil type, and the optimal phase Φ2opt and optimal gain G2opt thusobtained may be stored in advance in the phase-amplitude LUT 40A bybeing associated with the imaging region and coil type. In this case,since the optimal phase Φ2opt and optimal gain G2opt with high accuracyare obtained by referring to the phase-amplitude LUT 40A in Step S1, thereadjustment procedures in Steps S2 to S6 (especially Steps S3 to S5)may be omitted.

In the example described above, RF transmit pulses are transmittedsimultaneously via two transmission channels, i.e., the first and secondtransmission channels ch1 and ch2, by changing the phase Φ2 or gain G2and then the phase Φ2 and gain G2 of the second channel ch2 arereadjusted based on multiple items of data obtained for different valuesof the phase  2 and gain G2. Even for a case in which a fitting curve ofan ellipse is found from the B1 map data of the head, and the optimalphase Φ2opt is determined based on the angle of the major axis of theellipse and the angle of the major axis of an ideal ellipse,transmission is performed simultaneously via the first and secondtransmission channels ch1 and ch2.

In contrast, with the method described below, the optimal phase Φ2optand optimal gain G2opt are determined from two sets of B1 map data,including B1 map data obtained by means of transmission via the firsttransmission channel ch1 alone and B1 map data obtained by means oftransmission via the second transmission channel ch2 alone. This methodwill be referred to hereinafter as an independent-channel transmissionmethod. The independent-channel transmission method can find the optimalphase Φ2opt and optimal gain G2opt directly from the two sets of B1 mapdata without varying the phase Φ2 and gain G2 in multiple steps.

FIG. 10 is a flowchart showing exemplary procedures for finding theoptimal phase Φ2opt and optimal gain G2opt using the independent-channeltransmission method. In FIG. 10, steps equivalent to those in FIG. 5 aredenoted by the same step numbers as the corresponding steps in FIG. 5,and description thereof will be omitted.

With the independent-channel transmission method, in Step S10, theoptimal phase Φ2opt and optimal gain G2opt of the second transmissionchannel ch2 are determined based on first B1 map data collected by meansof transmission via the first transmission channel ch1 alone and secondB1 map data collected by means of transmission via the secondtransmission channel ch2 alone.

FIG. 11A shows an image which results when the first B1 map dataacquired using a thin phantom is displayed with plural contours, wherethe first B1 map data has been collected by means of transmission viathe first transmission channel ch1 alone with the second transmissionchannel ch2 turned off. On the other hand, FIG. 11B shows an image whichresults when the second B1 map data acquired using a thin phantom isdisplayed with plural contours, where the second B1 map data has beencollected by means of transmission via the second transmission channelch2 alone with the first transmission channel ch1 turned off.

Out of arbitrary lines passing through a center of the image, a linewhich maximizes uniformity of amplitude distribution will be referred tohereinafter as a maximum uniformity line. White arrows shown in FIGS.11A and 11B indicate the maximum uniformity lines of first and second B1map data, respectively.

When the phantom is spherical, ideally the maximum uniformity line ofthe first B1 map data collected by means of transmission via the firsttransmission channel ch1 alone corresponds in orientation to a magneticfield formed in the horizontal direction by the first transmissionchannel ch1, and thus has a horizontal orientation, i.e., an angle of0°. On the other hand, the maximum uniformity line of the second B1 mapdata collected by means of transmission via the second transmissionchannel ch2 alone corresponds in orientation to a magnetic field formedin the vertical direction by the second transmission channel ch2, andthus has a vertical orientation, i.e., an angle of 90°.

Although shape of the B1 map data varies depending on shape of thephantom, the above-described feature regarding the orientation of themaximum uniformity line is generally maintained.

In FIG. 11A, the angle of the maximum uniformity line is approximately0°, which is close to an ideal state. On the other hand, in FIG. 118,the angle of the maximum uniformity line is approximately 30°, whichdeviates approximately 60° from the ideal angle of 90°. An amount ofcorrection for use to correct the phase Φ2 of the second transmissionchannel ch2 can be determined in such a way that the angle of themaximum uniformity line for the second B1 map data will be 90°.Specifically, a difference between an angle (30° in this case) formed bythe respective maximum uniformity lines of the first and second B1 mapdata and an angle of 90° formed by the two maximum uniformity lines inideal conditions can be used as an amount of correction for use todetermine the optimal phase Φ2opt of the second transmission channelch2. In the example of FIGS. 11A and 11B, the amount of correction tothe phase Φ2 of the second transmission channel ch2 is 60°.

FIG. 12A shows an image resulting from a B1 map data collected by meansof simultaneous transmission via the first and second transmissionchannels ch1 and ch2 with the phase Φ2 of the second transmissionchannel ch2 kept at the initial value without making the correctiondescribed above. On the other hand, FIG. 12B shows an image resultingfrom a B1 map data collected by means of simultaneous transmission viathe first and second transmission channels ch1 and ch2 after correcting(readjusting) the phase Φ2 of the second transmission channel ch2 by 60°from the initial value. It can be seen from FIG. 12B that the correctionto (readjustment of) the phase Φ2 of the second transmission channel ch2has resulted in improvement of the B1 map data.

The maximum uniformity line can be extracted, for example, as follows.First, a center of the imaging area is found using arbitrary image datacollected in advance, such as sensitivity map data of receiver coils orpositioning (locator) image data. Next, in the B1 map data, straightlines are drawn radially from the center of the imaging area at desiredintervals such as 5-degree intervals and a representative value (e.g.,average value) of B1 values on each straight line is calculated. Theaverage value of B1 values can serve as an index which represents theuniformity of B1 values on the straight line. Thus, the straight linewhich maximizes the representative value (e.g., average value) of B1values can be extracted as the maximum uniformity line.

On the other hand, the amount of correction for use to readjust the gainG2 of the second transmission channel ch2 can be determined based on thefirst and second B1 map data collected by means of separate transmissionvia the first and second transmission channels ch1 and ch2, as with theamount of correction to the phase Φ2.

Specifically, after finding contribution of the RF transmit pulsetransmitted via the first transmission channel ch1 to the B1 uniformityand contribution of the RF transmit pulse transmitted via the secondtransmission channel ch2 to the B1 uniformity, the amount of correctionto the gain G2 is determined such that the RF transmit pulse transmittedvia the channel which makes greater contribution will have a largeramplitude.

For example, it can be seen from FIG. 11A that strength of the B1 formednear a center of imaging by the RF transmit pulse transmitted via thefirst transmission channel ch1 is low. On the other hand, it can be seenfrom FIG. 11B that high-strength B1 is formed near the center of imagingby the RF transmit pulse transmitted via the second transmission channelch2. Thus, it is desirable, in terms of increasing the uniformity, torelatively increase the gain G2 of the second transmission channel ch2than to increase the gain G1 of the first transmission channel ch1.

The contribution to the B1 uniformity can be found using any of variousindices. For example, by drawing straight lines radially from the centerof the imaging area at desired intervals such as 5-degree intervals inthe B1 map data, a representative value such as the average value ortotal value of B1 values on each straight line can be calculated. Then,a sum of the representative values of B1 values on the straight linescan be used as an index of contribution to the B1 uniformity.

In this case, the amount of correction Δgain2 to the initial value ofthe gain G2 can be calculated, for example, based on a ratio betweencontributions of the two transmission channels ch1 and ch2 to the B1uniformity as shown in Equation (3).

Δgain2 (dB)=20 log₁₀(SumB1ch2/SumB1ch1)  (3)

where SumB1ch1 is the sum total (antilogarithm) of B1 representativevalues on the straight lines on the B1 map data for the firsttransmission channel ch1 and SumB1ch2 is the sum total (antilogarithm)of B1 representative values on the straight lines on the B1 map data forthe second transmission channel ch2.

FIG. 23A shows an image resulting from B1 map data collected by means ofsimultaneous transmission via the first and second transmission channelsch1 and ch2 with the amount of correction Δgain2 to the initial value ofthe gain G2 of the second transmission channel ch2 set at −5 dB. On theother hand, FIG. 13B shows an image resulting from B1 map data collectedby means of simultaneous transmission via the first and secondtransmission channels ch1 and ch2 with the amount of correction Δgain2to the initial value of the gain G2 of the second transmission channelch2 set at +5 dB. In both FIGS. 13A and 13B, the amount of correctionΔgain2 to the phase Φ2 of the second transmission channel ch2 is fixedat 60°.

It can be seen from FIG. 13A that since the first transmission channelch1 whose contribution to the B1 uniformity is small has its gain G1 setrelatively high, the B1 strength in the center of imaging is low,reducing the B1 uniformity. On the other hand, it can be seen from FIG.13B that since the second transmission channel ch2 whose contribution tothe B1 uniformity is large has its gain G2 set relatively high, the B1strength in the center of imaging is high, improving the B1 uniformity.

Although in the example described above, both phase Φ2 and gain G2 ofthe second channel ch2 are readjusted (corrected) based on two sets ofB1 map data (i.e., the first and second B1 map data), the readjustmentof one of the phase Φ2 and gain G2 may be omitted.

Also, the user may be allowed to manually readjust one or both of thephase Φ2 and gain G2 by operating the input device 33 with reference toimages displayed on the display device 34 based on the two sets of B1map data.

Once the optimal phase Φ2opt and optimal gain G2opt are determined inthe manner described above, the imaging condition setting unit 40supplies various parameters and the like related to the optimal phaseΦ2opt, optimal gain G2opt, and pulse sequence to the sequence controllercontrol unit 41. This makes it possible to perform imaging scans. Theimaging scans are performed as follows.

As a scan start command is given to the sequence controller control unit41 from the input device 33, the sequence controller control unit 41supplies various imaging parameters acquired from the imaging conditionsetting unit 40 to the sequence controller 31. Based on the variousimaging parameters received from the sequence controller control unit41, the sequence controller 31 drives the gradient power supply 27,transmitter 29, and receiver 30, causing a gradient magnetic field to beformed in the imaging area in which the patient P is placed, and an RFsignal to be generated from the RF coil 24.

Consequently, an NMR signal generated by nuclear magnetic resonance inthe patient F is received by the RF coil 24 and supplied to the receiver30. Upon receiving the NMR signal from the RF coil 24, the receiver 30performs required signal processing, and then generates raw data by A/Dconverting the NMR signal into digital data. The receiver 30 suppliesthe generated raw data to the sequence controller 31. The sequencecontroller 31 supplies the raw data to the sequence controller controlunit 41, which then supplies the raw data as k-space data to the dataprocessing unit 42.

Next, the data processing unit 42 applies an image reconstructionprocess to the k-space data, thereby generating image data, appliesnecessary image processing to the image data, and displays the resultingimage data on the display device 34. This allows the user to check adiagnostic image.

That is, the magnetic resonance imaging apparatus 20 described above isconfigured to be able to adjust the phase and amplitude of the RFtransmit pulse appropriately according to imaging conditions such as theimaging region and/or RF coil type as well as readjust the phase andamplitude of the RF transmit pulse as required based on collected dataso as to make transmitted B1 more uniform. Thus, the magnetic resonanceimaging apparatus 20 can acquire more uniform transmitted B1 accordingto imaging conditions such as the imaging region and/or RF coil type.

While certain embodiments have been described, these embodiments havebeen presented by way of example only, and are not intended to limit thescope of the invention. Indeed, the novel apparatuses and unitsdescribed herein may be embodied in a variety of other forms;furthermore, various omissions, substitutions and changes in the form ofthe apparatuses and units described herein may be made without departingfrom the spirit of the invention. The accompanying claims and theirequivalents are intended to cover such forms or modifications as wouldfall within the scope and spirit of the invention.

1. A magnetic resonance imaging apparatus comprising: an imaging unitconfigured to carry out magnetic resonance imaging of a patient using atransmitting QD coil which allows at least one of phase and amplitude ofa radio-frequency transmit pulse on at least one input channel of thetransmitting QD coil to be adjusted independently of each other; and anadjustment unit configured to adjust at least one of the phase and theamplitude of the radio-frequency transmit pulse according to imagingconditions.
 2. The magnetic resonance imaging apparatus according toclaim 1, wherein: the imaging conditions include at least one of animaging region of the patient and the type of the QD coil; the magneticresonance imaging apparatus further comprises a storage unit configuredto store the imaging conditions and at least one of the phase and theamplitude of the radio-frequency transmit pulse, the at least one of thephase and the amplitude being associated with the imaging conditions;the adjustment unit acquires at least one of the phase and the amplitudeof the radio-frequency transmit pulse which meets the set imagingconditions from the storage unit as an initial value and adjusts the atleast one of the phase and the amplitude using the initial value of theacquired at least one of the phase and the amplitude.
 3. The magneticresonance imaging apparatus according to claim 2, wherein the adjustmentunit collects data while changing at least one of the phase and theamplitude from the initial value and adjusts the at least one of thephase and the amplitude based on the collected data.
 4. The magneticresonance imaging apparatus according to claim 3, wherein the adjustmentunit obtains an index which represents non-uniformity of aradio-frequency magnetic field formed by the QD coil, from the collecteddata and adjusts the at least one of the phase and the amplitude so asto minimize the index.
 5. The magnetic resonance imaging apparatusaccording to claim 4, wherein the data is adjustment data collectedusing a pulse sequence of a spin echo method with a flip angle setsmaller than a typical flip angle.
 6. The magnetic resonance imagingapparatus according to claim 4, wherein the data is a radio-frequencymagnetic field map.
 7. The magnetic resonance imaging apparatusaccording to claim 5, wherein the adjustment unit applies a process forenhancing strength variation to the adjustment data.
 8. The magneticresonance imaging apparatus according to claim 6, wherein the adjustmentunit applies a process for enhancing strength variation to theradio-frequency magnetic field map.
 9. The magnetic resonance imagingapparatus according to claim 2, wherein the adjustment unit acquires anaxial section image of the patient, determines an angle differencebetween a specific axis in the axial section image and an actual axis ofthe patient corresponding to the specific axis as a correction value,and corrects the initial value of the phase stored in the storage unitby the correction value.
 10. The magnetic resonance imaging apparatusaccording to claim 3, wherein the adjustment unit adjusts the at leastone of the phase and the amplitude so as to minimize forward power of arefocusing pulse measured while changing the at least one of the phaseand the amplitude.
 11. The magnetic resonance imaging apparatusaccording to claim 2, wherein: the QD coil is supplied with theradio-frequency transmit pulse from a first channel and a secondchannel; the adjustment unit adjusts at least one of the phase and theamplitude based on a first radio-frequency magnetic field map collectedby means of transmission via the first channel alone and a secondradio-frequency magnetic field map collected by means of transmissionvia the second channel alone.
 12. The magnetic resonance imagingapparatus according to claim 11, wherein the adjustment unit adjusts thephase based on an angle formed by maximum uniformity lines extractedfrom the first radio-frequency magnetic field map and the secondradio-frequency magnetic field map.
 13. The magnetic resonance imagingapparatus according to claim 11, wherein the adjustment unit adjusts theamplitude based on contribution to uniformity of a radio-frequencymagnetic field of the QD coil, the uniformity having been calculatedfrom the first and second radio-frequency magnetic field maps.
 14. Themagnetic resonance imaging apparatus according to claim 3, wherein theadjustment unit adjusts an output level of the radio-frequency transmitpulse based on at least one of the phase and amplitude subjected toadjustment.
 15. The magnetic resonance imaging apparatus according toclaim 14, wherein the adjustment unit adjusts the output level so as tominimize an index which represents non-uniformity of a radio-frequencymagnetic field collected by changing the output level.
 16. The magneticresonance imaging apparatus according to claim 14, wherein theadjustment unit adjusts the output level so as to maximize strength of asignal collected by changing the output level.
 17. The magneticresonance imaging apparatus according to claim 11, wherein theadjustment unit adjusts an output level of the radio-frequency transmitpulse based on at least one of the phase and amplitude subjected toadjustment.
 18. The magnetic resonance imaging apparatus according toclaim 17, wherein the adjustment unit adjusts the output level so as tominimize an index which represents non-uniformity of a radio-frequencymagnetic field collected while changing the output level.
 19. Themagnetic resonance imaging apparatus according to claim 17, wherein theadjustment unit adjusts the output level so as to maximize strength of asignal collected while changing the output level.
 20. The magneticresonance imaging apparatus according to claim 2, wherein the adjustmentunit acquires at least one of the phase and the amplitude of theradio-frequency transmit pulse which meets the imaging conditions fromthe storage unit, and multiplies the acquired at least one of the phaseand the amplitude by −1 when body position of the patient indicates alateral recumbent position.
 21. The magnetic resonance imaging apparatusaccording to claim 1, wherein the adjustment unit adjusts the amplitudeafter adjusting the phase.
 22. The magnetic resonance imaging apparatusaccording to claim 1, wherein: the imaging conditions include at leastone of an imaging region of the patient and the type of the QD coil; themagnetic resonance imaging apparatus further comprises a storage unitconfigured to store the imaging conditions and at least one of the phaseand the amplitude of the radio-frequency transmit pulse, the at leastone of the phase and the amplitude being associated with the imagingconditions; and the adjustment unit reads at least one of the phase andthe amplitude of the radio-frequency transmit pulse which meets the setimaging conditions from the storage unit and sets the read at least oneof the phase and the amplitude on a transmission channel of the QD coil,and thereby adjusts the at least one of the phase and the amplitude. 23.The magnetic resonance imaging apparatus according to claim 22, whereinthe at least one of the phase and the amplitude stored in the storageunit is determined so as to minimize an index which representsnon-uniformity of a radio-frequency magnetic field formed by the QDcoil, the index being obtained from data collected while changing atleast one of the phase and the amplitude.
 24. The magnetic resonanceimaging apparatus according to claim 22, wherein: the QD coil issupplied with the radio-frequency transmit pulse from a first channeland a second channel; and the at least one of the phase and theamplitude stored in the storage unit is determined based on a firstradio-frequency magnetic field map collected by means of transmissionvia the first channel alone and a second radio-frequency magnetic fieldmap collected by means of transmission via the second channel alone.